I. Field of the Invention
The present invention relates to medical imaging systems; more particularly, the present invention relates to a high density, highly integrated, APD photosensor (Avalanche Photodiode) array for scintillation crystal readout in detectors utilized in imaging systems using positron emission tomography (PET).
II. Background Information
Nuclear medicine is a unique medical specialty wherein radiation is used to acquire images, which show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions, which emanate from the body and are captured by a scintillation crystal, with which the photons interact to produce flashes of light or “events.” Events are detected by an array of photo detectors, such as photomultiplier tubes, and their spatial locations or positions are calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body.
One particular nuclear medicine imaging technique is known as Positron Emission Tomography, or PET. PET is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. Measurement of the tissue concentration of a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from positron annihilation. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by an annihilation event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Annihilation events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors, i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line of response, or LOR, along which the annihilation event has occurred.
An example of a PET method and apparatus is described in U.S. Pat. No. 6,858,847, which patent is incorporated herein by reference in its entirety. After being sorted into parallel projections, the LORs defined by the coincidence events are used to reconstruct a three-dimensional distribution of the positron-emitting radionuclide within the patient. PET is particularly useful in obtaining images that reveal bioprocesses, e.g. the functioning of bodily organs such as the heart, brain, lungs, etc. and bodily tissues and structures such as the circulatory system.
In order to minimize patient exposure to radiation, detectors utilized in PET imaging systems must be able to detect low levels of incident optical photons or ionizing particles. In such imaging devices it is often advantageous to employ radiation detection devices having internal gain; avalanche photodiodes (APDs) are commonly used in such devices to provide the desired detection sensitivity. An APD is a semiconductor device that is biased near the breakdown region such that charge generated as a result of the absorption of an incident photon is amplified in the APD itself as a result of a cascading effect as charge is accelerated by the high bias potential applied across the p-n junction of the device. In such imaging devices, it is desirable that the APD exhibit low noise and high gain. Certain devices, such as medical imagers (e.g., using gamma radiation), also require relatively large arrays (e.g., about 5 cm.sup.2 or larger) of high quality, low noise APDs.
One detector configuration in PET imaging systems utilizing APD arrays is configured in a one-to-one coupling configuration, where one APD is coupled to one scintillation crystal. To collect the maximum amount of light from the scintillator, the APD has to have the same surface area as the scintillator crystal to which it is coupled. This can result in large surface area APDs, which increases the APD noise and capacitance. The noise and capacitance of an APD is directly proportional to its surface area. As the noise and capacitance of the APD increases, its capacity to accurately determine the proper energy and timing of an event decreases. This results in poor PET detector performance and is a main reason why APDs are not commonly used as photosensors in detectors in PET imaging systems. There is a need for a photosensor array configuration that can take advantage of the high gain resulting from the use of APDs in the array while simultaneously reducing the noise and capacitance resulting from increasing the size of APDs when used in photosensors.
One detector which utilizes the concept of a high density photosensor readout coupled to a single scintillator is the SiPM (Silicon Photomultiplier). A SiPM uses a very dense array of Geiger-mode APDs which are typically resistively connected in parallel to provide a single readout channel. A significant drawback of this type of this type of detector is that each of the SiPM APD cells is nonlinear since they operate in Geiger mode. Detectors utilizing SiPM Geiger mode APDs in each cell operate in binary mode. Accordingly, the respective output of SiPM APD cells can only be zero or one. This is a fundamental problem, because each SiPM APD cell is only capable of counting one photon and cannot indicate that more than one photon has been received. For example, if two photons reach the same SiPM APD cell at the same approximate time, there is no way of knowing that two photons have been received by the SiPM APD cell. To increase the linearity of a SiPM, the SiPM cell density must be increased. However, increasing the cell density of a SiPM causes a decrease in the fill factor of the device. For most SiPM devices, there is a trade-off between linearity and fill factor. There is a need for a configuration of APD arrays that overcomes this limitation wherein an array of APD cells operate in proportional mode and are linear.
Some developers in the industry have investigated operating SiPMs below the breakdown voltage in proportional mode for CT imaging to overcome the inherent nonlinearity of the device. However, the proposed detector still uses a common readout of the summed SiPM cells. There is a need for a system that facilitates individual cell readout of the cells comprising an APD array.